1. Field of the Invention
This invention relates generally to the field of mechanical cardiac pumping devices, and, more particularly, to a ventricular assist device (VAD) and a total artificial heart (TAH) device and method of using same. More specifically, this invention relates to a VAD and a TAH that are optimized by the new method to produce customized pulsatile blood flow mimicking that of the healthy native heart for each individual patient case.
2. Description of Related Art
Introduction:
Some medical studies indicate: a) 400,000 new cases of congestive heart failure are diagnosed annually in the United States; b) a mortality rate of 75 percent in men and 62 percent in women; c) standard medical therapies benefit only a limited percentage of patients with ventricular dysfunction; and d) from 17,000 to 66,000 patients per year, in the United States alone, may benefit from a permanent implantable blood pump. Presently, potential cardiac transplant recipients with hemodynamic compromise (inadequate perfusion of the systemic circulation by the native heart) sometimes receive temporary mechanical circulatory support as a xe2x80x9cbridgexe2x80x9d to permit them to survive until cardiac transplantation is possible. It is foreseen that some day mechanical blood pumps will provide a cost-effective alternative to either cardiac transplantation or long term medical management of patients. It is to this end that the devices and methods described herein have been developed.
It is to be understood that for purposes of this document a xe2x80x9cventricular-assist device (VAD)xe2x80x9d is a mechanical blood pump that assists a diseased native heart to circulate blood in the body, and a xe2x80x9ctotal artificial heart (TAH)xe2x80x9d is another type of mechanical blood pump that replaces the native heart and provides all of the blood pumping action in the body.
In order for a VAD to function optimally, it must both complement the diseased native heart and make the combined output of the VAD and native diseased heart emulate the pumping action of the natural healthy human heart. That is, it should provide pulsatile flow similar to that of the healthy heart. In order for a TAH to function optimally, it must mimic the pulsatile pumping action of the natural healthy human heart. In either case, the device must be sized such that it fits within the required areas in the patient""s body. In order to minimize the size of the power supply portion of the device, each device (VAD or TAH) must use as little energy and as little power as possible to accomplish the required function. Thus, there is a need for bio-emulating efficient pump (BEEP) systems for VAD and TAH applications.
It is known that VADs can be implanted to assist a functioning heart that does not have adequate pumping capability. Often, however, residual cardiac function is not taken into account in the design of such devices, resulting in less than optimal effects. What is needed is a bio-emulating efficient pump (BEEP) system, which works in concert with the native human heart. The new VAD device and system and optimization procedure described herein utilize patient specific information concerning residual cardiac output to optimize the pumping action provided for each individual patient, thereby providing such a BEEP system. The TAH device and optimization procedure described in this document optimize the pumping function provided for each individual patient, thereby providing such a BEEP system which is customized for each such patient. Known Heart Pump Devices:
Previously, a number of devices were developed for blood pumping. Highly specialized pumps have been used to completely replace a biological heart which has been surgically removed. Such known heart pumps may be temporary, or permanently implantable. Temporary heart pump devices usually involve either: 1) an attempt to augment a compromised native heart while it recovers from surgery or some other short-term problem; or 2) use of the device as a xe2x80x9cbridgexe2x80x9d to extend the life of a patient by temporarily replacing the native heart until a suitable donor heart can be found for cardiac transplantation.
Many types of permanently implantable heart pumps have been proposed and several have been developed. Because the left ventricle of the heart, which pumps blood to the entire body except for the lungs, becomes diseased far more commonly than the right ventricle (which pumps blood only to the lungs), most heart pumps have been developed to assist or replace the left ventricle. Fewer pumps have been proposed, tested, and used for bi-ventricular function (i.e. assisting both the left and right ventricles).
Known mechanical blood pumps can be roughly divided into three major categories: a. pulsatile sacks; b. reciprocating piston-type pumps; and c. pumps with axial or centrifugal impellers. Each category has distinct advantages and disadvantages.
a. Pulsatile Sacks
Pulsatile sack devices are the most widely tested and used implantable blood pumps. These devices employ flexible sacks or diaphragms which are compressed and released in a periodic manner to cause pulsatile flow of blood. Sack or diaphragm pumps are subject to fatigue failure of compliant elements. They are generally used as temporary heart-assist devices, and they are mechanically and functionally different from the present invention described hereafter.
The intra-aortic balloon (IAB) counter-pulsation device, a pulsatile sack device, is readily available. It is a catheter-mounted intra-vascular device designed to improve the balance between myocardial oxygen supply and demand. The first successful clinical application of the balloon was reported by Kantrowitz et al. in 1968. The IAB is positioned in the thoracic aorta and set to inflate at the dicrotic notch of the atrial pressure waveform when monitoring aortic pressure. The diastolic rise in aortic pressure augments coronary blood flow and myocardial oxygen supply. The IAB is deflated during the isovolumetric phase of left ventricular contraction. The reduction in the afterload component of cardiac work decreases peak left ventricular pressure and myocardial oxygen consumption. These units are not portable and are limited to in-hospital critical care use only. Use of the IAB is now a standard form of therapy for a variety of patients with cardiovascular disease, primarily reserved for patients with deteriorating heart function while awaiting revascularization procedure. In 1993, nearly 100,000 IABs were inserted in the United States alone.
Another example of a pulsatile sack device is the Abiomed(trademark) BVS(copyright) device (Abiomed, Inc., Boston, Mass.). It is an externally placed dual-chamber device that is capable of providing short term univentricular or biventricular support. It has pneumatically driven polyurethane blood sacks and it is not intended for long-term support. Also, U.S. Pat. No. 4,888,011 to Kung and Singh discloses a hydraulically driven dual-sack system; and U.S. Pat. No. 5,743,845 to Runge discloses a sack-operated bi-ventricular assist device that balances the flow in the left and right side of the circulatory system.
b. Reciprocating Piston-Type Pumps
Several types of implantable blood pumps containing a piston-like member have been proposed to provide a mechanical device for augmenting or totally replacing the blood pumping action of a damaged or diseased heart. For example, the HeartMate(copyright) (Thermo Cardiosystems, Inc., Woburn, Mass.) is a pneumatically powered device that is implanted in the left upper quadrant of the abdomen. A pneumatic air hose exits from the lower half of the abdominal wall and is attached to a pneumatic power unit. Blood from the cannulated left ventricular apex empties into a pump, at which point an external control system triggers pumping. The blood chamber is pressurized by a pusher plate forcing a flexible plastic diaphragm upward. This motion propels the blood through an outflow conduit grafted into the aorta, the main artery supplying the body with blood. This device is unique in that the textured, blood-containing surface promotes the formation of a stable neointima, hence full anticoagulation is not necessary, only anti-platelet agents are required. This device is designed for left ventricular support only. It uses trileaflet polyurethane valves. There is an electrically powered version with percutaneous electric leads connecting the pump to external batteries.
The Thoratec(copyright) VAD (Thoratec Laboratories, Pleasanton, Calif.) is a pneumatically powered device that is placed externally on the anterior abdominal wall. Cannulas pass through the chest wall in a manner similar to that of a conventional chest tube. The device takes blood from the left ventricular apex and returns it to the aorta. Full systemic anticoagulation is required with this device. It can be used to support either ventricle and uses tilting disc type mechanical valves.
Novacor(copyright) (Cedex, France) produces an electrically driven device that is implanted in the left upper quadrant of the abdomen and the electric line and vent tube are passed through the lower anterior abdominal wall. This system also incorporates a polyurethane blood sac that is compressed by dual symmetrically opposed pusher plates. Blood is taken from the left ventricular apex and returned to the aorta. Full anticoagulation is required.
U.S. Pat. No. 3,842,440 to Karlson discloses an implantable linear motor prosthetic heart and control system containing a pump with a piston-like member which reciprocates in a magnetic field. The piston includes a compressible chamber in the prosthetic heart which communicates with the vein or aorta.
U.S. Pat. Nos. 3,911,897 and 3,911,898 to Leachman, Jr. disclose heart assist devices controlled in the normal mode of operation to copulsate and counterpulsate with the heart, respectively, and produce a blood flow waveform corresponding to the blood flow waveform of the assisted heart. The heart assist device is a pump connected serially between the discharge of a heart ventricle and the vascular system. This pump has cylindrical inlet and discharge pumping chambers of the same diameter and a reciprocating piston in one chamber fixedly connected with a reciprocating piston of the other chamber.
U.S. Pat. No. 4,102,610 to Taboada et al. discloses a magnetically operated constant volume reciprocating pump which can be used as a surgically implantable heart pump or assist. The reciprocating member is a piston carrying a check valve positioned in a cylinder.
U.S. Pat. Nos. 4,210,409 and 4,375,941 to Child disclose a pump used to assist the pumping action of the heart with a piston movable in a cylindrical casing in response to magnetic forces. A tilting-disk type check valve carried by the piston provides for flow of fluid into the cylindrical casing and restricts reverse flow.
U.S. Pat. No. 4,965,864 to Roth discloses a linear motor using multiple coils and a reciprocating element containing permanent magnets, driven by microprocessor-controlled power semiconductors. A plurality of permanent magnets is mounted on the reciprocating member. U.S. Pat. No. 4,541,787 to DeLong describes a pump configuration wherein a piston containing a permanent magnet is driven in a reciprocating fashion along the length of a cylinder by energizing a sequence of coils positioned around the outside of the cylinder.
U.S. Pat. No. 4,610,658 to Buchwald et al. discloses an implantable fluid displacement peritoneovenous shunt system. The device is a magnetically driven pump, which can be a reciprocating diaphragm, or piston type, or rotary pump.
U.S. Pat. No. 5,089,017 to Young et al. discloses a drive system for artificial hearts and left ventricular assist devices comprising one or more implantable pumps driven by external electromagnets. The pump utilizes working fluid, such as sulfur hexafluoride to apply pneumatic pressure to increase blood pressure and flow rate.
Larson et al. in a series of patents (1997-1999, U.S. Pat. Nos. 5,879,375; 5,843,129; 5,758,666; 5,722,930; 5,722,429; 5,702,430; 5,693,091; 5,676,651; 5,676,162) describe a piston-type pump for ventricular assist or total replacement, and associated driving equipment and power supply. The piston is an artificial heart valve, with valves that have at least two leaflets, acting as a check valve and reciprocating in a cylinder. The walls of the cylinder are a few millimeters thick because they contain the coils of a linear electric motor that must provide pumping power to the VAD. Around the artificial heart valve and inside the cylinder is a hollow cylindrical rare-earth permanent magnet, which is driven by the linear electric motor. In one embodiment one device is implanted in series to the aorta (left VAD), or another device is implanted in series to the pulmonary artery (right VAD), or two devices are used on both aorta and pulmonary artery (BI-VAD). In a second embodiment one device replaces the left ventricle, or another device replaces the right ventricle, or two devices replace the whole heart.
Measurements on experimental devices made with hollow pump cores indicate that such devices are too large to fit in the available space in the chest cavity in the aorta or pulmonary artery, due to the size of the coils necessary to drive the device. For a given volume of blood pumped per stroke, if the length of the cylinder is restricted such that the device fits lengthwise in the human body, then the diameter must be increased until the desired volume is reached. The outer diameter of the device is severely restricted by the surrounding tissue, and this leaves little room available in the diameter for the linear magnet motor. In a bi-ventricular application, if the axes of the two cylinders are located in parallel, then even more space is needed due to the diameters required; and if they are not parallel the magnetic fields of the two motors introduce additional electromagnetic losses because the linear magnet motors are not parallel. Even if the volumetric displacement of the device is reduced in order to fit in the available space at the expense of throughput, much of the outside diameter of the device must still be devoted to the linear motor. However, the most important disadvantage is that the linear motor is driving an annular magnet containing a one-way valve, so that the ferromagnetic material can not be in the core (center) of the motor coils, leading to lower efficiency.
At the geometric center (axis) of the motor described by Larson et al. is the artificial valve acting as the piston, and the blood itself. This structure introduces electromagnetic losses in the device that make it less desirable than devices that have ferromagnetic material in the geometric center (axis) of the motor coils. In addition, voltage propagates at constant velocity from coil to coil in the linear magnet motor of the Larson et al. device, and motion of the magnet carrying the artificial heart valve is coupled to this application of voltage, so that the application of current in the Larson et al. device is not optimized to minimize the power required to effect the blood-pumping action.
c. Pumps with Axial or Centrifugal Impellers
After pulsatile devices, rotary pumps, having either centrifugal or axial impellers, are the most widely used and tested devices. In centrifugal pumps, the blood flow enters axially into a centrifugal impeller, centrifugal acceleration increases the blood flow velocity, the flow exits radially, and the flow is subsequently decelerated to increase blood static pressure in the diffusion process. Most such centrifugal pumps provide continuous (non-pulsatile) flow; or flow with a small fluctuating pressure trace superimposed on a larger steady-pressure component, such as U.S. Pat. No. 5,928,131 to Prem and U.S. Pat. No. 6,179,773 to Prem and Kolenik.
Axial pumps direct blood flow along a cylindrical axis, which is in a straight (or nearly straight) line with the direction of the inflow and outflow. The impeller looks like an axial fan, or propeller, inside a nozzle. The impeller imparts acceleration to the fluid, and the subsequent deceleration (diffusion) process increases the blood pressure. Most such axial pumps provide continuous (non-pulsatile) flow.
Some types of axial rotary pumps use impeller blades mounted on a center axle, which is mounted inside a tubular conduit. As the blade assembly spins, it functions like a fan or an outboard motor propeller. Another type of axial blood pump, called the xe2x80x9chaemopumpxe2x80x9d uses a screw-type impeller with a classic screw (also called an Archimedes screw; also called a helifoil, due to its helical shape and thin cross-section). In screw-type axial pumps, the screw spins at very high speed (up to about 10,000 rpm). The entire haemopump unit is usually less than one centimeter (approximately 0.4 inches) in diameter. The pump can be passed through a peripheral artery into the aorta, through the aortic valve, and into the left ventricle. An external motor and drive unit powers it.
Axial and centrifugal pumps provide mostly steady (continuous) flow with an imperceptible high-frequency low-amplitude pulsatile component. Various mechanisms have been proposed to convert this practically steady-flow output into pulsatile flow. However, both axial and centrifugal impeller pumps introduce rapid acceleration and deceleration forces and large shear stresses in the blood. As is well known to those with ordinary skill in the art (Balje, 1981), both types of turbomachines (axial and centrifugal) are a balanced compromise between diameter and speed to provide the specified flow rate and pressure increase. Imposing limits in diameter in order to reduce shear stresses means that the optimum machine requires a higher-speed axial component. Imposing speed limits in order to reduce shear stresses means that the optimum machine requires a higher-diameter centrifugal component. It is well know to those with ordinary skill in the art (Wilson and Korakianitis, 1998) that small impellers that can fit inside the spaces available in the human body will result in high blood shear, due to the high operational speed required.
The Jarvik 2000(copyright) (registered trademark of R. Jarvik, New York, N.Y.) System consists of a small axial flow pump (about the size of a C-cell battery) that is placed in the left ventricular apex and pumps blood into the aorta. It is still currently being developed and will use external batteries and control electronics utilizing induction coils to carry the control signals through the skin. Power is also delivered transcutaneously.
Medical Complications:
According to several medical studies, the above devices are subject to a number of complications. Insertion of a cannula to feed a pump can cause damage to the left ventricle. At least 50 percent of patients who are supported for prolonged periods develop infections, including those associated with pneumatic lines or electrical leads. Septic emboli may occur, and the mortality rate is up to 50 percent. VADs may also activate the coagulation cascade, resulting in thrombi formation. This occurs in the approximate range of nine to forty-four percent of patients. Stasis of blood within the pump may lead to thrombus deposition. Right ventricular failure may occur peri-operatively with placement of a left VAD. The right heart failure rate may be as high as 33 percent, with one-fifth of those patients dying from the complication. Rapid recognition of this complication and implantation of a right VAD may reduce the mortality rate resulting from right heart failure. Hemorrhage occurs in about 27 to 87 percent of patients who require mechanical ventricular assistance. Hemorrhage is also related to inflow and outflow cannulae and to anticoagulation required with the devices.
One of the most important problems in axial and centrifugal rotary pumps involves the interface between the edges of the blades and the blood flow. The outer edges of the blades move at high speeds and generate high levels of shear. Red blood cells are particularly susceptible to shear stress damage, as their cell membranes do not include a reinforcing cytoskeleton to maintain cell shape. Lysis of red blood cells can result in the release of cell contents and trigger subsequent platelet aggregation. Lysis of white blood cells and platelets also occurs upon application of high shear stress. Even sublytic shear stress leads to cellular alterations and direct activation and aggregation of platelets. Rotary pumps generally are not well tolerated by patients for prolonged periods. In medical tests, animals placed on these units for a substantial length of time often suffer from strokes, renal failure, and other organ dysfunction.
The device and method of optimization disclosed herein minimizes the above, and other, known complications resulting from implantation of either a VAD or a TAH.
Desirable Pump Characteristics:
In many patients with end stage heart disease, there is enough residual function left in the native heart to sustain life in a sedentary fashion, but insufficient reserve for even minimal activity, such as walking a short distance. This residual function of the diseased native heart is typically not considered in the design of most VADs. Most known VADs are designed to assume complete circulatory responsibility and to receive blood from the cannulated ventricular apex of the particular ventricle they are xe2x80x9cassisting,xe2x80x9d in what is commonly called xe2x80x9cfill to emptyxe2x80x9d mode. It generally takes one or more contractions of the diseased native ventricle to supply enough blood to the VAD. Once a pre-specified volume of blood is accumulated in the VAD, then the ejection phase of the VAD is initiated. Thus, most known VADs operate in this xe2x80x9cfill-to-emptyxe2x80x9d mode that is in random association with native heart contraction, and can be installed in parallel to the native ventricle or in series. These constructions are not considered to xe2x80x9ccomplementxe2x80x9d the native heart, as does the present invention.
At least some residual cardiac function is present in the majority of patients who would be candidates for mechanical circulatory assistance. It is preferable for the natural heart to continue contributing to the cardiac output even after a mechanical circulatory device is installed. This points away from the use of total cardiac replacements and suggests the use of assist devices whenever possible. However, the use of assist devices also poses a very difficult problem. In patients suffering from severe heart disease, temporary or intermittent crises often require artificial pumps to provide bridging support which is sufficient to entirely replace ventricular pumping capacity for limited periods of time. Such requirements arise in the hours or days following a heart attack or cardiac arrest, or during periods of certain life threatening arrhythmias. Therefore, there is an important need to provide a pump and method that can meet a wide spectrum of requirements by providing two different and distinct pumping functions, assisting the native heart and total substitute pump support.
The present invention provides a cardiac ventricular-assist device and method of optimizing any design of VAD or TAH wherein the amount of power required by the device is minimized to the extent necessary to complement the cardiac output of the native heart, and no more. In this manner, the weight and size of the device are kept within suitable reasonable ranges to permit placement of the VAD/TAH within the body of the subject patient using the new device.
The present invention further provides a VAD and method wherein the principles of unsteady thermodynamics and fluid mechanics are used to provide a uniquely optimized pulsatile blood flow which complements the cardiac output of the individual native human heart. It is to be understood that throughout this document, when the terms xe2x80x9coptimizexe2x80x9d and xe2x80x9ccomplementxe2x80x9d are used in reference to the devices and systems of the present invention, it is meant that at each heart beat and stroke of the VAD (used here to mean either the L-VAD, R-VAD, BI-VAD or TAH as described below), several actions are carefully timed such that:
a) the native heart is allowed to pump as much blood as it can on its own before the VAD is activated;
b) as the blood-ejection phase of the native heart nears completion, the VAD is energized to provide additional pumping action;
c) the additional pumping action reduces the back pressure in that native ventricle so that the native ventricle pumps more than it would have pumped unaided;
d) the timing of the action, length of pumping stroke, and rate of pumping (stroke displacement versus time and resulting power input versus time) of the VAD are related to the native heart ejected blood volume and rhythm in a manner that minimizes power input to the VAD while meeting physiological constraints;
e) the optimization processes in d) take into account the dynamic interaction between the native heart and the VAD; and
f) the optimization process and the control scheme are integrated with the resulting changes in blood ejected per heart beat and heart rate (beats per minute) by the combined action of the native heart and the VAD.
Before turning to the Figures, it is considered useful to provide some introductory material. The present invention, described below, is distinct from each of the three categories of mechanical circulatory support devices previously described, and consolidates the advantages and avoids the disadvantages of each category. First, it is carefully noted that several of the devices described in the known art mention that the power input is xe2x80x9coptimizedxe2x80x9d, but they do not describe how this is accomplished. The optimization method described herein can be applied to all existing VAD and TAH devise that have been devised to date, or will be devised in the future.
The pump of the present invention has ferromagnetic material as the solid center of the motor coils, thus providing a more compact arrangement of the electromagnetic fluxes than pumps with non-ferromagnetic centers, and simultaneously permitting reduction of electromagnetic losses in use. Ultimately this permits placement of a device that can pump sufficient volume per stroke at the outlet of the native ventricles and allows the power supply to be smaller than was possible with previous cardiac pumping devices. The remote hydraulic drive and power supply/controller assembly are located in the abdomen, thus allowing practically all available space in the vicinity of the heart for use by the device. Power is transmitted hydraulically from the abdomen to the blood pump in the vicinity of the heart. Also, electromagnetic losses are not introduced by the location of the two pumping devices (artificial heart valves) in non-parallel configuration in the vicinity of the aorta and pulmonary artery.
Details of the dynamics of the pumping action of the human heart have been incorporated for the first time into the design of the VADs and TAHs in the present invention. Understanding these details:
1) is essential for optimization of the timing of unsteady-flow events in the heart-pumping cycle;
2) directly impacts the optimum geometric shape of the artificial devices; and
3) identifies prerequisite means to minimize shear stresses on the blood (reducing blood-cell lysis) and optimizing energy flows (reducing the power input required to produce the required blood flow and pressure characteristics).
The adult heart is located between the lungs and is about the size of a large grapefruit, weighing 0.2 to 0.5 kg (0.44 to 1.1 pounds), depending on the size of the individual. The cardiovascular system performs two major tasks: it delivers oxygen and nutrients to body organs; and removes waste products of metabolism from tissue cells. Its major components are: the heart (a two-sided biological pump); and the circulatory system of elastic blood vessels (veins and arteries) that transport blood. As an example, the heart of a 70 kg (154 pounds) human circulates about 6 kg (13.2 pounds), or 6 L (6.34 qt.s), of blood.
The human heart is divided into four chambers: the right atrium and right ventricle; and the left atrium and left ventricle. The walls of the chambers are made of a special muscle called the myocardium that contracts rhythmically under electric stimulation. The left and right atria are separated from each other by the atrial septum; and the left and right ventricles are separated from each other by the ventricular septum.
In the circulatory system, blood returns by the venous system from the body and enters the heart through the right atrium, then subsequently blood enters the right ventricle. Each time the right ventricle contracts, it propels this blood (low in oxygen content) into the lungs, where it is enriched with oxygen. Pulmonary veins return the blood to the left atrium, then subsequently the blood enters the left ventricle. The left ventricle, which traditionally has been considered as the main pumping instrument of the heart, ejects the blood through the main artery, the aorta, to supply oxygenated blood to the various organs of the body. The organs use the oxygen and with capillary action between the arterioles and the venules return the blood to the venous system and the right atrium. The pumping action of the left and right side of the heart generates pulsatile flow and pressure on the aorta and pulmonary artery, discussed further below.
Blood is kept flowing in this pulsatile cycle by a system of four one-way valves in the heart, each closing an inlet or outlet in one of the heart""s four chambers at the appropriate time in the cardiac cycle. The valve system helps maintain a pressure difference between the right and left sides of the heart. The aortic valve and the pulmonary valve each have three tissue cusps (leaflet flaps), referred to as xe2x80x9csemilunar valvesxe2x80x9d because of the crescent shape of these cusps. The tricuspid and mitral valves separate the atria from the ventricles. The mitral valve has two cusps and the tricuspid valve has three cusps. In addition, the cusps have thin chords of fibrous tissue (chordae tendineae), which tether the valves to the ventricular walls. When the ventricles contract, small muscles in their walls (papillary muscles) restrict closure of the mitral and tricuspid valve leaflets, preventing them from overextending.
Electric currents control the pumping motion of the heart. The currents originate in the sinus node (the heart""s natural pacemaker), a microscopic bundle of specialized cells located in the superior portion of the atria. The currents travel through a network of conducting fibers to the atrioventricular or AV node, the bundle of His, and the Purkinje fibers. The electric currents cause impulses that are transmitted and propagate in a wave fashion through the muscle fibers of the left and right atria to the atrioventricular node, located on the juncture between the right and left sides of the heart where the right atrium and right ventricle meet. From the atrioventricular node, they travel along the bundle of His and the Purkinje fibers through the muscles of the right and left ventricles. Most currents in the heart are less than a millionth of an Ampere, but they exert a powerful influence on the heart muscle.
The new VAD utilizes electromagnetic coils to drive a high-ferromagnetic-constant driving magnet in a reciprocating fashion so as to act as a piston for hydraulic fluid. The resultant movement of hydraulic fluid through the system in turn moves another magnet, which is annular, and which also drives in a reciprocating fashion. The movement of the driven annular magnet in turn moves still another magnet, an annular valve seat magnet, which supports a one-way valve. This valve seat magnet is located inside the annular driven magnet, the two magnets sharing a common center axis, hence coupling them together. The one-way valve pushes blood through the ascending aorta of the heart when the valve is pushed forward, and allows blood to flow freely past when the one-way valve is moved backward.
The present invention provides a ventricular-assist device and method for optimizing same that can be utilized to assist either the left ventricle (L-VAD) or right ventricle (R-VAD) of the native human heart or, if necessary, to assist both cardiac ventricles (BI-VAD). The L-VAD, R-VAD and BI-VAD devices all utilize principles of unsteady fluid mechanics to provide a uniquely individualized optimized pulsatile blood flow for each particular patient.
In an alternative embodiment, a total artificial heart (TAH) device that utilizes the principles of unsteady fluid mechanics provides a uniquely individualized optimized pulsatile blood flow for each particular patient. The optimized pulsatile blood flow mimics that of the native heart while simultaneously minimizing the power required to drive the TAH device.
Accordingly, it is among the goals of the present invention to provide a cardiac pump (VAD or TAH) device and system, and method for controlling and operating same which permit customized, optimized xe2x80x9cassistxe2x80x9d or xe2x80x9ctotalxe2x80x9d (xe2x80x9ccompletexe2x80x9d) cardiac pumping support for an indefinite period of time. Under appropriate conditions, the new VAD acts synergistically with the native heart to provide a seamless augmentation to the otherwise suboptimal output of the diseased native heart. This allows the new pump device (VAD) to take advantage of the natural, non-hemolytic pumping action of the native heart to the fullest extent possible to minimize red blood cell lysis, and to reduce mechanical stress on the VAD system pump, requiring less volume, less energy, and hence allowing longer pump life and longer battery life.
Accordingly, in furtherance of the above objects and goods, the present invention is, briefly, a method of optimizing a mechanical cardiac pumping device includes modeling the physical system, or portions thereof, of the patient who will receive the mechanical cardiac pumping device and identifying an operating condition of the native heart to which the device will respond. The model is used to determine the required blood volume to be ejected from the device and an initial estimate of the power required to be provided to the mechanical cardiac pumping device is provided in order to provide the required ejected blood volume. The resultant ejected blood volume is evaluated with data obtained from the model and the estimate of the power requirement is then updated. The above steps are iteratively performed until the power required to obtain the necessary ejected blood volume is identified. Possible variations of power and pumping rate that allow the mechanical cardiac pumping device to provide the required volume are determined and the variation that best matches the physiological constraints of the patient and minimizes the power required by the mechanical cardiac pumping device is selected. The steps are iteratively performed until the mechanical cardiac pumping device is optimized to respond to each desired operating condition of the native heart.
The mechanical system for accomplishing the new method is, briefly, a system for assisting cardiac ventricular function, the system including a hydraulic pumping assembly and a cardiac ventricular assist device (VAD) in fluid communication with the hydraulic pumping assembly, wherein the hydraulic pumping assembly includes an encapsulated hydraulic pump having a pumping chamber for retaining hydraulic fluid therein. The pumping chamber has opposed first and second ends and at least one electromagnetic coil surrounding the pumping chamber. A substantially solid high ferromagnetic-constant magnet is disposed longitudinally, slideably and reciprocally within the pumping chamber to act as a piston for driving hydraulic fluid within the pumping chamber in response to signals from a battery/controller assembly. A fluid line has a first end and a second end. The first end of the fluid line is connected to and in fluid communication with the first end of the pumping chamber and the second end of the fluid line is connected to and in fluid communication with the second end of the pumping chamber. The VAD is in fluid communication with the fluid line at a point on the fluid line after the point of connection of the check valve and before the connection of the second end of the fluid line and the second end pump chamber. A battery/controller assembly is operatively connected to the check valve and to the at least one electromagnetic coil to provide electric power and control signals to the pump. The battery controller assembly is in electrical communication with the native heart of the patient using the system, to thereby receive signals corresponding to physiological parameters from the native heart for transfer to the VAD.
The new VAD device is, briefly, a device to assist the function of a cardiac ventricle, the device having a first magnet with an open center and formed of high ferromagnetic-constant material. A first vessel of the device surrounds the first magnet and defines a space in fluid communication with the blood flow output great vessel associated with the diseased ventricle of a patient using the device, the first magnet being movable within the first vessel in substantially fluid-tight relation thereto. A second magnet is formed of high ferromagnetic-constant material in magnetic communication with the first magnet, so that the magnetic fluxes of the first magnet and the second magnet affect each other, so that the first magnet and the second magnet are biased toward and tend to lock to one another, to thereby move in the same direction as one another. A second vessel encases the second magnet and defines a space and is movable within the space in substantially fluid-tight relation to the second vessel, the space being defined by the second vessel being in fluid communication with a hydraulic pump for actuation the second magnet. A one-way valve is connected to the first magnet, the one-way vale being movable with the first magnet, and adapted to cause movement of blood from the diseased ventricle to and into the great vessel associated with the diseased ventricle.
These and other advantageous features of the present invention will be in part apparent and in part pointed out herein below.